This type of radiation detector is used in a medical diagnostic apparatus such as a nuclear medical diagnostic apparatus (ECT: Emission Computed Tomography), e.g. a PET (Positron Emission Tomography) apparatus or a SPECT (Single Photon Emission Computed Tomography) apparatus, for detecting (simultaneously measuring) radiation (e.g. gamma rays) released from radioisotopes (RI) introduced into a patient and accumulated in a site of interest, and obtaining sectional images of RI distribution in the site of interest.
A PET apparatus will be described by way of example. The PET apparatus detects, with opposite gamma-ray detectors, two gamma rays released in directions of about 180° apart from a site of interest of a patient, and is constructed to reconstruct a sectional image of the patient when these gamma rays are detected simultaneously (counted simultaneously). The gamma-ray detectors used in the PET apparatus for simultaneously counting the gamma rays may have scintillators that emit light in response to incident gamma rays released from the patient, and light-sensitive elements (e.g. photomultiplier tubes) for converting the light emitted from the scintillators to electric signals.
FIG. 10 is an outline view of a conventional example. A radiation detector 110 includes a two-stage scintillator block 101 having a scintillator array of two-stage structure, for example. This two-stage scintillator block 101 is formed of a scintillator array upper part 111F and a scintillator array lower part 111R. The scintillator array upper part 111F and scintillator array lower part 111R are manufactured separately, and have an adhesive layer 102 interposed to bond the two parts finally. Thus, the radiation detector 110 includes the scintillator array upper part 111F and scintillator array lower part 111R, a light guide 120 optically coupled to the two-stage scintillator block 101, and four photomultiplier tubes 131, 132, 133 and 134 optically coupled to this light guide 120.
The scintillator array upper part 111F and scintillator array lower part 111R have two-dimensional close arrangements of scintillators 101SF and scintillators 101SR divided by light reflex materials 112 inserted in between. In the embodiment described hereinafter and in FIG. 10, the scintillators are arranged three-dimensionally; eight in X-direction, eight in Y-direction, and two stages in Z-direction, and thus a total of 128 (=8×8×2). The light guide 120 has a light guide lattice frame (not shown) which is a combination into a lattice form of strips (not shown) formed of optical elements such as light reflex materials. This light guide lattice frame defines numerous cubicles.
A specific method of manufacturing the two-stage scintillator block 101 is as follows. (1) First, when manufacturing the scintillator array upper part 111F, a lattice frame is prepared by combining a plurality of plate-like optical elements conforming to the height (the length in the direction of depth of gamma ray incidence) of the scintillators 101SF. (2) Before storing the lattice frame in a receptacle that can store the lattice frame, a transparent optical binding material is poured into that receptacle. (3) The scintillators 101SF are stored after storing the lattice frame in the receptacle, and the optical binding material is allowed to harden in this state. (4) The scintillator array upper part 111F having the hardened optical binding material, lattice frame and scintillators integrated together is taken out of the receptacle, and its outer shape is adjusted, whereby the scintillator array upper part 111F is manufactured. (5) The scintillator array lower part 111R also is manufacture by the same method as (1)-(4), and the two parts are bonded by the adhesive layer 102.
The principle of detection by the two-stage scintillator block 101 will now be described with reference to FIGS. 11 and 12. FIGS. 11 and 12 are explanatory views concerning identification of radiation detection in conventional examples. In FIGS. 11 and 12, reference RI denotes a radiation source, reference W denotes spaces (pitch) between the scintillators, and references L1 and L2 denote parallax errors. In principle, gamma rays released from positions distant from a visual field center often obliquely enter the scintillators of radiation detectors (radiation detectors D3, D4 in FIG. 11 and radiation detectors MD3, MD4 in FIG. 12).
As shown in FIG. 11, radiation detectors D having scintillators not divided in the direction of depth of gamma ray incidence not only detect a correct position but also detect an incorrect position (see the shaded portions in FIG. 11). That is, the view error becomes large gradually from the visual field center toward peripheries, and sectional images obtained with the PET apparatus are inaccurate.
On the other hand, as shown in FIG. 12, radiation detectors MD having scintillators divided in the direction of depth of gamma ray incidence provide the following functions and effects. That is, attenuation time of the emission pulse of light produced from incident gamma rays will be described with reference to radiation detectors MD having scintillators divided such that the scintillator array with the shorter attenuation time (scintillator array upper part 111F in FIG. 10) lies on the gamma ray incidence side, and the scintillator array with the longer attenuation time (scintillator array lower part 111R in FIG. 10) on the photomultiplier tube side (i.e. the side reverse of the gamma ray incidence side). In the case of these radiation detectors MD, improvement is sought to detect positions of emitted gamma rays accurately (see the shaded portions in FIG. 12) and obtain sectional images with increased accuracy even when gamma rays obliquely enter the scintillators of the radiation detectors MD (see Patent Documents 1 and 2, for example).
The gamma ray positions of the scintillator array with the short attenuation time and the scintillator array with the long attenuation time, stacked in the direction of depth of gamma ray incidence are specifically detected and identified as follows. As shown in FIG. 13, use is made of analog signals SF (signals of the scintillator array with the short attenuation time) or analog signals SR (signals of the scintillator array with the long attenuation time) which are electric signals outputted from the photomultiplier tubes which are the light-sensitive elements. Integration values of digital signals are calculated as shown in FIG. 14.
In FIG. 14, T1 refers to an intermediate time at an intermediate point from emission start to emission end of emission pulse produced in the scintillator block, AT1 refers to an intermediate additional value which is a sum of digital signals A from the emission start to the intermediate time T1, BT1 refers to an intermediate additional value which is a sum of digital signals B from the emission start to the intermediate time T1, T2 refers to the emission end, AT2 refers to a total additional value which is a sum of digital signals A from the emission start to the emission end T2, and BT2 refers to a total additional value which is a sum of digital signals B from the emission start to the emission end T2. In FIG. 14, reference A denotes an integration value of digital signals resulting from A/D conversion of analog signals SF (signals of the scintillator array with the short attenuation time), and reference B denotes an integration value of digital signals resulting from A/D conversion of analog signals SR (signals of the scintillator array with the long attenuation time).
The PET apparatus has an A/D converter, an adding device, a threshold calculating device, a mean value calculating device and a discriminating device (none being shown). The A/D converter converts the analog signals SF or analog signals SR shown in FIG. 13 into digital signals. The adding device successively adds the digital signals converted by the A/D converter. The addition by the adding device obtains the above intermediate additional value AT1 or intermediate additional value BT1, and the total additional value AT2 or total additional value BT2, respectively. The threshold calculating device calculates values AT1/AT2 or BT1/BT2 which is the intermediate additional value AT1 or intermediate additional value BT1 divided by the total additional value AT2 or total additional value BT2. This AT1/AT2 or BT1/BT2 is shown as threshold value. The mean value calculating device calculates mean value K from a maximum and a minimum of the threshold values calculated by the threshold calculating device. The discriminating device discriminates whether the threshold values calculated by the threshold calculating device are larger or smaller than the mean value K, thereby to detect and identify a gamma ray position.
[Patent Document 1]
Unexamined Patent Publication H6-337289 (pages 2-3, FIG. 1)
[Patent Document 2]
Unexamined Patent Publication No. 2000-56023 (pages 2-3, FIG. 1)